Optical Coherence Tomography (OCT) is an emerging non-invasive biomedical imaging technology that can perform cross-sectional imaging of tissue microstructures in vivo and in real-time. OCT is analogous to ultrasound, except that it uses low coherence light, rather than acoustic waves. The echo delay time or the depth of light backscattered from the tissue is measured using a technique referred to as low coherence interferometry.
OCT has significant advantages over other medical imaging technologies. Medical ultrasound, magnetic resonance imaging (MRI), and confocal microscopy are ill suited to morphological tissue imaging, as ultrasound and MRI have insufficient resolution for imaging microstructures, while confocal microscopy lacks the ability to image deeply enough (i.e., beyond several hundred micrometers in highly scattering tissues), which is required for morphological tissue imaging.
As indicated above, a fundamental aspect of OCT is the use of low coherence interferometry. In conventional laser interferometry, the interference of light occurs over a distance of meters. In OCT, the use of broadband light sources (i.e., light sources that can emit light over a broad range of frequencies) enables the interference to be generated within a distance of micrometers. Such broadband light sources include super luminescent diodes (i.e., super bright light emitting diodes (LEDs)) and extremely short pulsed lasers (i.e., femto second lasers). White light can also be used as a broadband source.
Essentially, the combination of backscattered light from the sample arm and reference light from the reference arm gives rise to an interference pattern, but only if light from both arms have traveled “substantially the same” optical distance (where “substantially the same” indicates a difference of less than a coherence length). By scanning the mirror in the reference arm, a reflectivity profile of the sample can be obtained. Areas of the sample that reflect back larger amounts of light will create greater interference than areas that reflect back smaller amounts of light. Any light that is outside the short coherence length will not generate any interference. This reflectivity profile, referred to as an A-scan, contains information about the spatial dimensions and location of structures within the sample. An OCT image (i.e., a cross-sectional tomograph generally referred to as a B-scan), may be achieved by laterally combining multiple adjacent axial scans at different transverse positions by utilizing a depth-priority imaging sequence (e.g., fast axial scanning followed by slow transverse scanning).
FIG. 1 (Prior Art) schematically illustrates a conventional OCT system. This system includes a Michelson interferometer that uses a low coherence light source 20. The light source is coupled to an OCT probe 24 in the sample arm and to a reference arm 28 through an optic fiber coupler or beam splitter 22. The sample arm delivers an optical beam from the light source to a target 26 (generally a tissue sample) and collects the backscattered light. The reference arm performs depth scanning by using a translating retro-reflective mirror or a phase-controlled scanning delay line (not separately shown). A backscattered intensity versus depth data set is developed with an axial scan. Two- or three-dimensional data sets formed by multiple adjacent axial scans are obtained by scanning the OCT beam along the transverse direction after each axial scan. A photodetector 30 produces a corresponding analog signal comprising the data set. The analog signal is processed by detection electronics module 32, which produces corresponding digital data. The resulting data set can be displayed using a computer 38, as a false-color or gray-scale map, to form a cross-sectional OCT image.
Unlike confocal microscopy, the transverse and axial resolutions of OCT are determined independently. The axial resolution Δz is based on the coherence length of the light source and is inversely proportional to the source spectrum bandwidth Δλ, according to the following relationship:
                              Δ          ⁢                                          ⁢          z                =                              (                                          2                ⁢                                                                  ⁢                ln                ⁢                                                                  ⁢                2                            π                        )                    ⁢                      (                                          λ                0                2                                            Δ                ⁢                                                                  ⁢                λ                                      )                                              (        1        )            where λ0 is the source center wavelength. The transverse resolution, Δx, is determined by the transverse focused spot size, in a manner similar to that in conventional microscopy, according to the following relationship:
                              Δ          ⁢                                          ⁢          x                =                                            2              ⁢              λ                        π                    /                      d                          2              ⁢              f                                                          (        2        )            where d is the beam spot size on the objective lens, and f is the focal length of the objective lens.
Implementing real-time OCT with continuous focus tracking in a depth-priority imaging sequence can be very challenging, since an extremely high tracking speed (on the order of a few meters/second) and a high repetition rate (in the kHz range) are required. Focus tracking is not critical when using a large transverse focused spot size. However, as a tighter focus (or higher transverse resolution) is utilized, the transverse resolution will deteriorate faster at depths farther from the focal plane. The depth of focus b (or the confocal parameter) reduces quadratically with the spot size diameter Δx according to the following relationship:
                    b        =                              π            ⁢                                                  ⁢            Δ            ⁢                                                  ⁢                          x              2                                            2            ⁢            λ                                              (        3        )            
For example, the depth of focus reduces from about 200 μm to about 50 μm when the transverse resolution increases from 10 to 5 μm. Conventional OCT has a relatively low transverse resolution, between about 20 μm and about 40 μm, and focus tracking is not generally necessary for low resolution OCT. However, low transverse resolution degrades image contrast. Even with coherence gating along the axial direction, photons that are backscattered within the focal spot size by different scatterers (e.g., by cells or cell organelles) will likely be simultaneously detected and averaged, causing loss of contrast. Therefore, a high transverse resolution is much preferred. To maintain a high transverse resolution at various depths, focus tracking is needed. As indicated in Equation (3), small changes in the spot size diameter result in large changes to the depth of focus. Thus, dynamic focus tracking becomes very important for maintaining the focused spot size throughout the entire imaging depth.
Conventional real-time OCT imaging is achieved by fast axial scanning followed by slow transverse scanning, and the image consists of multiple adjacent axial scans at different transverse locations. A 2-3 mm axial scan generally takes less than 0.5 milliseconds during real-time imaging, requiring focus tracking at a velocity of about 4-6 meters per second, which is extremely difficult to achieve in a compact scanning device. FIG. 2 illustrates the rapid depth scanning of tissue 42 by an incident beam 40, and the relatively slow transverse scans that are used. In this conventional technique for OCT scanning, focus tracking means that the focus point is rapidly tracked at each different transverse location before moving to the next transverse location, which is very challenging to achieve.
One simple non-dynamic focus tracking approach is to acquire a sequence of images with the focus gradually shifted into the sample, and then to fuse together the in-focus image zones through post-image processing. Unfortunately, precise image registration is difficult to achieve, and the effective frame rate is reduced by at least the number of focal zones taken to generate one “in-focus” image. In contrast, dynamic focus tracking seeks to simultaneously track the imaging beam focus and the coherence gate throughout the entire imaging depth, by maintaining a preferably zero or near zero optical path length (OPL) difference between the reference and sample arms, as determined from the focal plane within the imaging depth.
It would be desirable to provide alternative techniques for implementing focus tracking in OCT imaging. It would be particularly desirable to provide a technique enabling a relatively fast frame rate, and which enables the beam spot size to be maintained over a variety of focal depths. Relatively fast frame rates will facilitate the use of OCT imaging with live tissue (i.e., live biological specimens). Preferably, the frame rate will be faster than the respiratory rate of the specimen, to avoid blurring due to respiratory activity.